CT scanners have become one of the most widespread modalities for diagnostic imaging because of the soft tissue contrast and spatial resolution of images that can be produced by the CT scanners. Recent studies indicate that CT imaging can greatly increase the detection of small non-calcified nodules as compared with planar chest radiography. Accordingly, there is an increasing interest in CT scanning. By way of example, lung cancer screening has been receiving significant attention in recent years.
Since human exposure to x-rays is generally harmful, the risk of patient radiation exposure should be significantly less than the risk of undetected pathologies that can be treated at early stages to affect a better outcome. Accordingly, the benefit of CT screening must outweigh the risk of the additional radiation exposure for a particular application. For CT to be a successful modality in screening, detectors and scanning technologies that can assure delivery of a very low dose of radiation to the patients without sacrificing imaging quality, should be developed.
X-ray detectors play a very important role in the performance of a CT scanner. CT detector technologies have continually changed over years in a tremendous effort to improve performance. The first clinical CT systems used detectors that were typically based on high-pressure inert gases (usually xenon). Examples of these models are the GE 7800, GE 8800, and GE 9800. A disadvantage of xenon detectors is their low detection efficiency due to the relatively low density of gas. Although various efforts have been made to improve the performance of these detectors by increasing the pressure of the xenon chamber, the gas detector still falls short in terms of detection efficiency as compared to solid-state detectors coupled to scintillators. An advantage of the xenon detector is its low cost.
To improve efficiency, newer models began to use solid state detectors based on silicon (Si) photodiodes optically coupled to scintillating materials such as CdWO4 or custom produced scintillating ceramics. Si detectors coupled to scintillators overcome the efficiency limit of xenon detectors. In these detectors, optical photons, induced by an x-ray interaction in the scintillator, travel toward the photodiodes and produce an electrical signal. A total light signal originates from multiple interactions of x-ray photons with the scintillator and is consequently converted to an electrical signal by the photodiodes and read out by external electronic circuitry. The analog signal from the detector then is digitized with data acquisition electronics.
Currently, CT systems use such a detector that includes Si photodiodes that are optically coupled to a scintillator. The need to obtain the soft tissue contrast and spatial resolution desired in reconstructed images places high demands in terms of intrinsic spatial resolution and dynamic range on the detector. To meet such demands, the typical size of pixels in CT detectors is about 1 mm, and as there can be a tremendous flux of x-rays delivered to the detectors (about 100×106 photons/mm2/second in air without the patient) a large dynamic range of linear response is required.
The detector is typically operated in an integrating mode where the detector generates a signal proportional to the total energy deposited as a function of time. An incident x-ray photon undergoes a photoelectric interaction with a scintillation converter which emits visible or ultra violet light. This light reaches the photodiode component which provides an electrical signal (current) proportional to the energy fluence of x-rays. Thus as the current is being monitored, these detectors are often referred to as operating in ‘current mode’.
Current mode read out integrates both the signal and noise from the detector and electronics over time. Furthermore, the x-rays in CT have a broad spectrum of energies. When either the count rate or the x-ray energies are low, the signal from x-rays must exceed a noise level produced by the detector and readout electronics. Thus there is a minimum threshold in terms of x-ray flux that can be reliably detected, which increases as the x-ray energy decreases. Thus, in conventional current mode CT detectors, there is a distinct non-zero lower limit on the dynamic range. The noise limit in clinical whole body CT scanners places a limit on soft tissue contrast. Despite these limitations, current mode x-ray detectors are used currently in virtually all clinical x-ray systems including CT and digital radiography.
Conventional CT systems utilize detectors that convert energy into current signals that are integrated over a time period, then measured and ultimately digitized. A drawback of such detectors however is in their inability to provide data or feedback as to the number and energy of photons detected. As previously mentioned, one problem with these detectors is that there is a lower limit of detection defined by noise in the detector such that a certain pixel with no incident radiation may produce some signal due to thermal and shot noise.
There are additional deficiencies of integrating systems, such as not taking advantage of statistical information carried by each photon (e.g., 3 photons of 30 keV carry the same information as one photon of 90 keV) and not using information about the energy of the counted photons. For example, because a poly-energetic x-ray spectrum is used, each detected photon also contributes different information to the resulting image depending on density and elemental composition of the examined tissue. Utilization of the energy information carried by individual photons can lead to further improvement of the quality of the image and/or reduction of the radiation dose. This is accomplished by optimal energy weighting to increase soft tissue contrast. Because of these limits inherent in current mode detectors, there has been a mounting effort to move away from this technique and explore photon counting detectors for x-ray imaging applications such as CT, digital radiography, and mammography.
However, a drawback of photon counting detectors is that these types of detectors generally cannot count at very high x-ray photon flux rates typically encountered with conventional CT systems. The very high x-ray photon flux may lead to nonlinear detector system responses and ultimately a saturation of the detector system. This saturation can occur particularly at a detector location wherein the boundary of the subject is imaged and a very small thickness of the tissue (or none) is placed between the detector and the x-ray generator. For example, when x-rays are passed through a person through a thicker part of the person's body, the x-rays have a higher probability of being attenuated. In other words, the flux of x-rays that passes through the person and thus reaches the detector, is much reduced at the thicker part of the person's body. However, when the beam of x-rays is passed through a thinner part of the person's body near the edges, a very high count rate is detected. In other words, a very high x-ray flux is detected at the edges of the person. This problem can be mitigated to a certain degree by placing a bow-tie shaped filter between the subject and the x-ray generator in order to equalize total attenuation and reduce x-ray photon flux at the subject peripheries. However, this technique has limitations due to problems optimizing the filter due to the fact that members of the subject population do not have a uniform body shape and are not exactly elliptical in shape.
A number of techniques have been tried to avoid saturation problems in CT detector systems. One such technique involves reducing the x-ray flux by using a lower current at the x-ray tube and hence a lower x-ray flux. This will lead to an increased scan time which will be needed to acquire images that meet the required statistics to obtain good image quality. The increased acquisition time in turn can lead to a blurring of images due to organ movement within the scan time. Other solutions include software correction algorithms. Unfortunately, software solutions may create image artifacts because of an inability to perfectly replace the saturated data. Another solution is explored in U.S. Pat. No. 6,953,935 B1 (GE patent).